In the past, in nuclear medicine diagnosis, nuclear medicine diagnosis apparatuses have been developed where a radioactive isotope (RI) is injected into a living body, and the concentration distribution of RI within the body is captured by a one-dimensional or two-dimensional detector to acquire useful diagnosis information such as a lesion, blood flow, fatty acid metabolism and the like.
As the nuclear medicine diagnosis apparatus, there are a SPECT apparatus (Single Photon Emission Computed Tomography) for injecting single photon nuclides into a test body, counting gamma rays emitted from the body, and reconstructing the cross-sectional concentration distribution, and a PET apparatus (Positron Emission Computed Tomography) for capturing a pair of gamma rays emitted from a positron nuclide in the 180-degree direction by their coincidence to count, and reconstructing the cross-sectional concentration distribution.
Conventionally, crystal scintillator detectors such as BGO, GSO, LSO and the like have generally been used as a gamma ray detector of the nuclear medicine diagnosis apparatus. The crystal scintillator detector has time and energy resolutions with high accuracy, but is limited in (reaction) position resolution due to the crystal size. Particularly, the position resolution in the depth direction (DOI: Depth-Of-Interaction in gamma ray emission direction) of the gamma ray reaction point is the order of centimeters. Therefore, the image quality deteriorates due to parallax, and to compensate for deterioration, measures are taken such as combining with a CT apparatus with high image quality and the like.
Further, as a detector with high position resolution of gamma ray interaction, there are semiconductor detectors such that many layers of silicon strip are stacked. To obtain sufficient radioactive stopping power in the semiconductor detector, it is necessary to stack at least hundred or more layers of silicon strip with a layer thickness of 0.5 mm. Therefore, with the apparatus increased in size, many semiconductor elements and a large number of readout channels are required and result in the problem of being expensive.
Meanwhile, it has been proposed applying a gamma-ray detector using liquid xenon (Xe), liquid krypton (Kr), or liquid argon (Ar) as a reaction medium of gamma ray to a PET apparatus (for example, see Patent Document 1). As shown in FIGS. 12(a) and 12(b), in a detector module 100 as described in Patent Document 1, a large number of photomultiplier tubes 102 are disposed on the sides and top of a rectangular chamber filled with liquid xenon 104, and electric field wires 106 are installed in the vertical direction along the side. Further, a collector pad 110 is installed on the bottom which is the incident plane of gamma ray, and a shutter system 112 is installed on the medium side of the collector pad 110.
In the detector module 100, when a gamma ray 150 enters the liquid xenon 104 from the collector pad 110 side, scintillation light is emitted from an interaction point P where the gamma ray 150 and liquid xenon 104 interact with each other, while xenon molecules are ionized, and ionization electrons are produced near the emission point. The scintillation light is detected by the photomultiplier tube 102, while the ionization electrons travel at a constant velocity toward the collector pad 110 within the liquid xenon 104 set in a drift electric field of 1 kV/cm. The ionization electrons drifting within the liquid xenon 104 enter the collector pad 110 via the shutter system 112, and the incident position is specified.
Then, signals output from a plurality of photomultiplier tubes 102 are analyzed, a reaction time point that the light is emitted and a first three-dimensional position of the interaction point P are determined, the position of the interaction point P is specified in two-dimension from the incident position of the ionization electrons entering the collector pad 110, and a second three-dimensional position is determined based on an arrival time point of the ionization electrons with respect to the time point that the light is detected.
[Patent Document 1] JP 2005-532567